The present invention can be understood more fully if image enhancement techniques developed for x-ray radiography (and in particular x-ray mammography) are detailed first.
Two aspects of traditional radiographic x-ray imaging, spectrum optimization and scatter rejection (or reduction), are of particular concern in specialized fields such as mammography and angiography. A variety of point, slit, and slot scanning systems have been designed to optimize the spectrum and scatter reduction, and special filters have been employed to shape the source spectrum for both mammography and angiography. See, R. Nelson, Dissertation, "High Resolution Slit Scan Techniques for Digital Radiography", Department of Radiological Sciences, University of California, Los Angeles (1986). For example, in the case of dual-energy k-edge subtraction angiography, the desire for a very narrow bandwidth intense x-ray beam resulted in the replacement of a conventional x-ray tube source with an approximation to a pulsed x-ray laser source: a synchrotron. E. Rubenstein, et al., SPIE vol. 314, Digital Radiography (Sept. 14-16, 1981). The use of a highly directional, narrow bandwidth, slit scanning beam from a synchrotron (a source of potentially very short pulses on the order of tens of picoseconds, see N. Schwentner, et al., Nuclear Instruments and Methods vol. 167, p.499-503 (1979)) presented opportunities to achieve additional scatter reduction.
Although it is difficult to exploit radiation beam properties such as polarization and coherence for scatter reduction at radiographic energies (typically above 15 Kev), it is possible to consider employing energy--selective, angle-selective, and time-of-flight (TOF) methods to limit large angle scatter and small angle scatter contributions (including contributions from multiple-scattered photons which emerge with a direction vector similar to that of the unscattered transmitted beam) to the image spatial contrast resolution and signal-to-noise ratio (SNR). The importance of utilizing all three methods in an imaging system has been understood by researchers in the field of positron emission tomography. See J. Llacer, et al., 1 IEEE Trans. Nucl. Sci. NS-20, p. 282-293 (1973). In addition, the principle of TOF has been exploited in the field of medical ultrasound for many years.
Mechanisms which can provide additional scatter rejection in a radiographic imaging system can be used either to enhance the image contrast resolution or alternatively to allow expansion of the beam size from a slit to a slot or area beam (thereby reducing image acquisition time) while maintaining an acceptable level of image contrast resolution. Unfortunately, implementing energy-selective x-ray spectroscopy with a detector array experiencing very high count rates is difficult and compromises may be required. See Nelson, et al., U.S. Pat. No. 4,937,453 (1990).
Implementing a TOF imaging system by gating or modulating an array of x-ray detectors which are designed to provide high contrast resolution and efficiency represents another challenging problem. Of course, additional scatter reduction is still possible if low-tech concepts are implemented. For example, an inexpensive spatial filter such as a focused x-ray grid could be added to remove some fraction of the large angle x-ray scatter in the plane of the slit.
Clinical x-ray mammography, which employs a film-screen receptor and a molybdenum-anode x-ray tube, is currently used as a mass-screening technique for breast disease. Two views of a breast are usually acquired in order to help separate overlapping objects, providing a very simple form of tomography. However, certain risks are associated with x-ray examinations since x-ray radiation is also ionizing, and, therefore, the exposure to which should be minimized. Minimizing exposure equates to limiting the frequency and number of exams and limiting exams to patients having a minimum recommended age. Conventional and unconventional (including CT and stereotactic) imaging techniques have been developed for x-ray mammography with the goals of improving image contrast resolution and improving detection of disease while lowering patient risk and exposure.
In x-ray mammography it is desirable to use a range of x-ray (electromagnetic) energies that will enhance the radiologists' ability to differentiate normal from diseased tissue while limiting patient radiation dose to tolerable levels. Unfortunately, in general, the use of such an x-ray energy range results in scattered x-ray photons comprising a significant fraction of the transmitted beam. Additional problems can arise due to the energy-dependent filtering action of a breast when a broad band mammography x-ray source is used (i.e. beam hardening). Both x-ray scatter and beam hardening problems can be reduced by compressing the breast to be imaged (which also helps reduce patient motion problems) while additional scatter reduction can be obtained by using an air gap and a focused x-ray grid (collimator). If the breast is sufficiently compressed, a focused grid need not be used at all. Both the air gap and the focused grid function as spatial filters, although the angular selectivity of the grid is relatively poor compared to what can be implemented in the optical realm (due in part to the energy range in which the grid is used and the angular distribution of the unfocused x-ray source). The size of the air gap cannot be too large due to magnification effects. The use of air gaps and grids represent conventional, widely practiced methods for improving image quality when a broad energy bandwidth relatively large area x-ray beam is used in clinical mammography.
The use of (preferably) narrow bandwidth, time-resolved or TOF imaging systems in x-ray mammography in a clinical setting is still impractical due to the lack of an affordable sufficiently brilliant pulsed source (synchrotrons are relatively expensive) and the lack of an inexpensive gated detector capable of high contrast resolution and efficiency (film-screen receptors are not sufficient and are still dominant in clinical mammography). Rationalizing the use of such expensive equipment to lower the cost of an expensive angiographic procedure is far easier than rationalizing the same for mammography, a mass-screening examination which is relatively inexpensive. If a TOF system could be implemented for x-ray mammography then a primary limitation on acquisition time (other than exposure to x-rays) is patient motion, i.e. the image or image segment must be acquired before patient motion affects the image and thereby becomes a problem. The integrated x-ray fluence requirements will be the same whether integration is over one intense short pulse or a rapid sequence of many weak short pulses. The power per pulse (for a single intense pulse or a rapid sequence of weak pulses) is not likely to be of concern at mammography x-ray energies since a tolerable fraction of unscattered photons are transmitted for the case of a typical compressed breast. Contrast this with TOF optical mammography imaging problems. In the case of TOF optical imaging attempts to image the unscattered component of the optical beam transmitted through a typical compressed breast, within the time constraints imposed by patient motion, could pose a radiation safety risk (e.g. burns).
Although a cost-effective, narrow bandwidth TOF x-ray mammography system is not available, it is still desirable to implement optimized beam filtration and scatter reduction techniques. A common approach to limiting transmitted scatter is to reduce the area of the x-ray beam. Given time constraints for image acquisition (patient motion and exposure) and limitations on the heat or electrical current capacity of the x-ray tube as well as the tube focal spot distribution, a slit or slot scanning format may be acceptable. Ideally, the size of the slit or slot would be appropriate for the thickness (and material composition) of the breast being imaged. The range of x-ray energies commonly used in film-screen mammography primarily exhibit two scattering mechanisms: Rayleigh or coherent scattering (elastic scattering without energy loss) and Compton scattering (inelastic scattering). See Radiologic Health Handbook p. 133 and 438, U.S. Department of Health, Education, and Welfare (January 1970); and see W. Veigele, Atomic Data Tables Vol. 5, p. 51-111 (1973) (for coherent and incoherent scatter cross-sections for hydrogen, carbon, nitrogen and oxygen see pages 66-69). For the range of energies used in x-ray mammography coherent scatter can comprise a significant fraction of the total scatter component of the transmitted x-ray beam reaching the detector. The affect on the resultant image appears to be a reduction of contrast and SNR. See P. Johns, et al., Med. Phys. 10(1), p. 40-50 (1983); and H. Barrett, et al., Radiological Imaging, vol. II, p. 631-635 (1981) (scatter point spread function). It should be noted that Rayleigh scattering is the primary concern when using visible and near-infrared wavelengths. See A. Ishimaru, Wave Propagation and Scattering in Random Media, vol. 1, p. 18 (1978). The Rayleigh or coherent component of x-ray scatter is typically small-angle scatter (although the coherent scatter peaks at an angle which appears to be much larger for x-ray mammography than for optical mammography).
An x-ray mammography TOF technique would ideally allow separation of unscattered x-ray photons from scattered x-ray photons (i.e. small angle scattered and large angle scattered (including multiple scattered x-ray photons which exit from the breast aligned with the unscattered beam)). Breast compression would still be very desirable because the contribution to the detected signal from the unscattered component is diminished exponentially with increasing breast thickness. A compressed breast or breast region also ensures that the tissue volume being imaged is of uniform thickness. A uniform thickness reduces the need for an extremely large detector dynamic range (which is a problem for mammography film-screen detection systems) and timing difficulties related to initiating gating in a TOF system. Image degradation from x-ray scatter remains a problem for clinical x-ray mammography imaging systems. Small angle scattered x-rays degrade the desired image far less than large angle scattered x-rays. In addition, the information content of small angle scattered x-rays could be potentially useful since these x-rays may sample a tissue volume which is similar to the desired tissue volume. The information content of small angle scatter x-ray can be enhanced if the tissue volume which can be sampled by all or part of the detected beam is restricted (for example by compressing the breast or limiting the size of the incident x-ray beam).
A perceived advantage of a TOF technique can be added to an x-ray mammography system design if large angle scatter (which is primarily Compton scatter) propagating approximately along the unscattered x-ray beam direction can be removed. This can be accomplished in part if a narrow bandwidth directional source is used, if a detection system offers energy-dependent directional discrimination capabilities, and if the beam-aligned Compton scattered photons have lost sufficient energy to be rejected by the detector. At mammography x-ray energies this condition is most likely to occur if photons undergo multiple Compton scattering. See H. Barrett, et al., Radiological Imaging vol. I, p. 321 (1981). The energy-dependent directional discrimination capability of the detector can also be used to remove part of the Compton or the Rayleigh scatter component which emerges from the breast with too large an angle with respect to the unscattered beam direction. See Nelson, et al., U.S. Pat. No. 4,958,368; and Nelson, et al., U.S. Pat. No. 4,969,175. If energy discrimination is not available, then spatially limiting the size of the x-ray beam becomes more important since this reduces the x-ray cross-talk contribution from adjacent tissue volumes to the unscattered x-ray photon beam exiting the breast.
Before leaving the topic of x-ray radiography it is useful to review additional x-ray imaging techniques capable of providing three dimensional information. In one case x-ray fluorescence was used to measure iodine concentration in thyroids. See H. Barrett, et al. Radiological Imaging vol. II, p. 661-662, (1981). Computed tomography (CT) enhances the available data set by acquiring projections from many directions, permitting three dimensional information to be synthesized from two dimensional data sets. Tomography, the poor-mans' CT, provides an image of a specific layer of a body part with (preferably) minimum distortion from the surrounding layers. Recently classical film-based tomography has been modified to include a digital acquisition system. Data from multiple angled projections can be combined to synthesize (by tomosynthesis) a three dimensional image. See H. Barret, et al., Radiological Imaging, Vol. 2, p. 368-370 (1981). Also see D. Nishimura, et al., SPIE vol. 314, Digital Radiography, p. 31-36 (1981); and J. Liu, et al., IEEE Trans Medical Imaging, Vol. 8, No. 2, p. 168-172, 1989). Although scatter reduction is considered desirable for typical transmission radiography applications, there are many techniques where x-ray scatter, including backscatter, has been used for imaging and densitometry. See J. Battista, et al., Phys. Med. Biol., vol. 22(2), p. 229-244 (1977); and (A. Jacobs, et al., SPIE vol. 206, p. 129-134 (1979) (backscatter imaging). Given the right conditions at least a fraction of the scattered x-rays can carry useful information.
The ability to limit the level of photon scatter and photon scatter distribution in a detected beam is important for optical mammography imaging as well as x-ray mammography. In optical mammography the maximum breast thickness for which a transmitted signal resulting from unscattered photons can be detected is dramatically less than for x-ray mammography. This implies that for a typical compressed breast there is little benefit gained from using optical photon pulses which are shorter than some minimum temporal width since practically all of the optical photons exiting along the desired direction are scattered. This does,not mean, however, that a TOF technique utilizing only unscattered photons could not be used for a sufficiently thin subject. See M. Duguay, et al., Applied Optics, vol. 10, No. 9, p. 2162-2170 (1971). The scale of this unscattered optical photon beam problem is readily apparent if one considers the work of early researchers concerned with optical propagation in blood, in particular the scatter cross-section and the mean cosine of the forward scatter component as a function of wavelength. See C. Johnson, IEEE Trans. Bio-Medical Engineering, vol. BME-17, No. 2, p. 129-133 (1970); and G. Pedersen, et al., Biophysical Journal, vol. 16, p. 199-207 (1976). As is the case for x-ray mammography, beyond some thickness of breast in order to generate a desired image (with adequate photon statistics) using optical photons with "optimal" information content the amount of energy absorbed by the breast being imaged becomes unacceptable. The imaging technique must be modified. In conventional x-ray mammography the x-ray tube potential might be increased and the x-ray receptor could be changed (e.g., Xeromammography). The compromise involves imaging with x-ray photons energies which will provide less than optimal information content (about absorption in tissue) in exchange for an acceptable patient dose.
For optical mammography imaging a modified TOF technique could be used to limit contributions from scattered photons which have sampled tissue outside the tissue volume of interest. The objective is to try to maximize the relevant information content of the imaged scattered photons which could help characterize the tissue volume which is being examined. See J. Maarek, et al., Med. & Bio. Eng'g & Com., p. 407-413 (July 1986). The drawback of allowing a longer flight time for photons is that the distribution of possible paths for the imaged photons and the acceptable angular distribution of the exiting imaged photons also increases. Unfortunately, in the optical case, the capability to filter photons which follow inappropriate paths is minimal (i.e., there is minimal spectral discrimination). Undesirable scatter contributions can be reduced by spatially limiting the size of the optical beam and using collimation which provides angular selectivity (which is also implemented in x-ray mammography).
Another problem for optical mammography imaging is that entrance and exit surfaces of the object to be imaged (i.e., skin surfaces) are typically not smooth. Optical photons can be greatly affected by the surface structure of the skin whereas x-ray photons are relatively insensitive to skin surface conditions. An optical coupling fluid or gel can help reduce the effects of an irregular skin surface. See U.S. Pat. Appln. Ser. No. 08/480,760, filed June 1995. Optical properties such as polarization (which is relatively impractical to utilize in x-ray mammography) may also be exploited. The ability to reduce degrading effects of optical scatter can be of benefit for other time-resolved techniques and is not limited to the TOF approach.
In recent years, broad beam light sources (sometimes referred to as "light torches") having a relatively wide spectral bandwidth in the visible and near-infrared range have been used for breast imaging. Broad beam light transmitted through a breast is typically recorded by a video camera and viewed on a video monitor or analyzed by a computer. However, the ability to discriminate between various types of tissue in a breast via this technique is reduced since the transmitted beam has a wide spectral bandwidth and the captured radiation is largely comprised of scattered radiation (i.e. contrast is lost). Light may be absorbed, transmitted, scattered, and reflected to different degrees by various types of tissue making it difficult to extract information about the nature of any tissue. Detection limits for this technique have generally been restricted to lesions which are no smaller than what a physician can detect by palpitation. Therefore, this technique is not particularly advantageous and has fallen out of favor in the United States.
As the present applicants described in now issued U.S. Pat. Nos. 4,649,275, 4,767,928, 4,829,184, 4,948,974, and pending patent application Ser. No. 08/480,760, filed Jun. 6, 1995, a collimated (i.e. focused) continuous wave (CW) or rapidly pulsed light (i.e. non-ionizing electromagnetic radiation including near-ultraviolet, visible, infrared, microwave, etc.) source of narrow spectral bandwidth (such as is generated by a filter lamp, LED, laser, a waveguide, a phased array, etc.) can be used to produce a beam or a number of beams of light having relatively small spatial dimensions appropriate for acquiring images of a breast with high spatial contrast resolution. The narrow spectral bandwidth of the beam, along with other beam parameters (such as polarization, directional qualities or angular distribution, etc.) enable improved characterization of the composition of the breast material being imaged. Additional information can be obtained by acquiring images at other wavelengths with narrow spectral bandwidths (and/or other modifications to beam parameters).
For transmission imaging it is preferred that the light source be positioned on one side of the breast and a receiver, such as a photodetector (radiation detector), be positioned on the opposite side to record transmitted light. For backscatter imaging the photodetector will be on the same side of the breast as the source. As is shown in FIGS. 1 and 2, in many instances it is preferred that the breast be compressed between compression plates, flattening (or shaping) the entrance and/or exit surface(s) while reducing the typical distance the radiation must travel before exiting the breast and being detected. An additional benefit is that the region being imaged tends to be of a more uniform thickness when compression plates are employed.
The type of optical (radiation) source (CW, modulated CW, pulsed, etc.), as well as other possible properties such as beam coherence, wavefront phase, polarization, angular and spectral distribution (the source may be tuneable), are altered by absorption and scattering as the beam propagates through the breast and plates. The radiation exiting the breast can be analyzed using various forms of external collimation such as air gaps, focused lenses (single lenses, lens combinations, holographic or diffractive lenses, etc.), narrow spectral bandwidth filters, directionally-sensitive filters, narrow spectral bandwidth and directionally-sensitive filters (for example, multilayer films, interferometers, etc.), polarized filters, amplifiers (including nonlinear amplifiers), optical shutters and mechanical apertures, holographic or diffractive spatial filters as well as acousto-optic devices which exhibit high angular sensitivity, fiber optics and amplified fiber optics, light pipes, waveguides, masks, focused arrays, etc.
Image resolution can be influenced by adjusting the cross-sectional area of the optical beam(s). The advantage of using a radiation beam of relatively small dimensions (either its two dimensional area cross-section which is typically defined as being normal to the beam axis or three dimensional volume cross-section in the case of a short pulse or a source with a short coherence length) is the ability to limit single and multiple scatter cross-talk contributions to the measured signal from neighboring tissue volumes. A number of factors such as the thickness of tissue, the uniformity of the region being imaged, the scatter reduction mechanisms employed, whether a time-resolved or diffusive-wave optical technique is implemented, etc., influence acceptable beam dimensions. For example, in a practical imaging system, there are almost no unscattered (ballistic) photons for thicknesses of breast tissue greater than a few centimeters. The ability to exploit minimally scattered photons (the snake component) and extend the acceptable range of tissue thickness is limited. The implication is that the utilization of breast compression is highly advantageous when making use of the snake component when imaging a typical breast. An additional drawback is the increased cross-talk within the beam relative to what can be achieved with a time-resolved technique (such as TOF) which successfully records only ballistic photons. Thus, if the snake component is utilized for optical imaging it is beneficial to use a smaller beam cross-section in comparison to the case in which optical imaging involves only the ballistic component.
The electromagnetic properties of various normal and diseased breast tissues may exhibit wavelength dependence. Thus, acquiring images at different wavelengths of light as well as examining the effects of tissue on other electromagnetic parameters (e.g., direction vector, polarization, phase, amplitude, temporal profile, coherence, etc.) may aid in distinguishing between various types of tissue. High resolution images may be obtained with a variety of scanning techniques: FIGS. 2a and 2b show a point beam or multiple point beam which can be used in a raster scan format. The transmitted light beam can be collimated by a variety of means. This approach can be extended to include a single line or multiple line scan format as shown in FIG. 2c. High speed two-dimensional imaging is shown in FIG. 2d. In this case collimation (such as fiber-optics or light pipes) can be introduced into one or both compression plates. FIG. 3 shows a (patterned) mask collimator which can be used to generate multiple beams. In all cases collimation may be used to produce a beam or beams of relatively small cross-section and directional nature. These attributes can be used to help exclude unwanted scatter in the detected beam.
If two or more sources providing light beams of differing wavelengths (i.e., .lambda..sub.1 and .lambda..sub.2) are spatially separated as shown in FIG. 1b, then narrow spectral bandwidth filters can be used between plate B and the detectors for each wavelength such that the detector for .lambda..sub.2 rejects light of wavelength .lambda..sub.1 which is scattered into the path of the .lambda..sub.2 beam. In this case the spectral filter functions as a collimator, rejecting a component of the transmitted beam which can only be attributed to scatter. By positioning source 1 (for .lambda..sub.1) adjacent to source 2 (for .lambda..sub.2) the scatter contribution from source 1 into itself (near the boundary with source 2) can be estimated by measuring the .lambda..sub.2 component at the location of source 1. This assumes that radiation of wavelengths .lambda..sub.1 and .lambda..sub.2 have similar scattering and absorption properties for the type of tissue being imaged. Another technique is to have sources 1 and 2 incident at the same location, but source 2 would be tilted with respect to source 1. If source 1 and source 2 have the same properties, then source 2 should be pulsed at a distinctly separate time relative to source 1 being pulsed (use a temporal offset) to help minimize beam cross-talk. The source 2 component measured at the location of the source 1 detector can be used to estimate a scatter correction in some instances. This measurement could, taking a different perspective, also be attributed to a new (virtual) beam which was created at a greater effective depth than source 1.
Thus, one approach to estimating scatter corrections or contributions is to use two sources with different wavelengths. Another technique is to use two sources with different polarizations (or simply alter the polarization of a single source) and acquire two measurements at different times. The source 1 detector can be used as a second detector or a separate detector can be employed. If a beam splitter can be used to separate the exiting beam (transmitted and/or backscattered) into two parts (or separate the exiting beam components directly) then two detectors can be used to make simultaneous measurements of the beam components.
Optical (non-ionizing radiation) tomography utilizing a collimator can be employed in a variety of fashions. For example, as shown in FIG. 6, an object, such as a breast 30 may be imaged by a source of radiation 32 generating a one or two dimensional radiation beam, a detector 34, and a collimator 36 disposed between the source 32 and the detector 34. In this way multiple two dimensional images may be obtained simultaneously, thereby providing a three dimensional image of the object. For example, as shown in FIG. 7, a line source 42 or linear array of point sources may irradiate the object to be scanned such as a breast 44. Transmitted radiation then passes through a collimator 46, and then is detected by a detector 48, such as a two dimensional array of detectors, or a camera. FIG. 11b shows an arrangement similar to that shown in FIG. 7 but the stationary collimator of FIG. 7 is replaced with a reciprocating collimator.
An optical structured (patterned) collimator (see FIGS. 3 and 4) such as a fiber optical bundle, mask or honeycomb-like device introduces its own transfer function into the transfer function of the imaging system (which includes the source and its collimator, the detector and its collimator, and the optical properties of the breast). Thus, the signal recorded by the detector(s) represents the superposition of all elements of the imaging chain. In addition, a fiber of the fiber bundle may be seen by more than one detector element. The adverse effects of an optical structured post-collimator pattern (such as a fiber bundle) on image quality can be reduced by moving the optical structured (patterned) collimator in a reciprocating fashion in front of the detector (See FIG. 8a, 11b).
Active collimation (in which the subject is modified rather than the optical beam in order to achieve scatter reduction) can be implemented by using compression plates to compress the region of the breast which is to be imaged. This reduces the effective volume of tissue a photon is likely to sample (and thus suffer additional scatter) before exiting the breast. The use of active collimation can be of particular value when time-resolved optical imaging techniques are employed. A variety of time-resolved optical imaging techniques (e.g. time-of-flight, holography, heterodyne, homodyne, Raman amplification, etc.) in development for use with highly scattering media exploit temporal or phase properties of the radiation field. See R. Berg, et al., SPIE Vol. 1511, p. 397-424, (1993). For example, if the light (radiation) source is pulsed and the pulse length is sufficiently short, time-of-flight (TOF) imaging and analysis (typically based on the "ballistic" and sometimes the "snake" component(s) of the radiation field) can be employed. Photon diffusive wave imaging (and spectroscopy) techniques (also referred to a frequency domain or photon migration or photon density wave techniques) may represent an alternative to time-resolved optical methods. Amplitude and phase modulation (and even phase encoding) have been utilized for composition and location identification. Photon diffusive wave computed tomography (CT) imaging has also been implemented. In addition, tomographic reconstruction techniques (several which make use of a reference medium) based on diffusion equation approximations to transport theory for thick tissue have been investigated. See, e.g., J. Fujimoto, 4 Optics & Photonics News 9-32 (1993); and Medical Optical Tomography, SPIE Vol. 1511, (1993).
In addition, it has been disclosed that the manner in which radiation interacts with a medium can be altered by the presence of an acoustic field. See, e.g., A. Korpel, Acousto-Optics (1988); and F. Marks, et al., "A Comprehensive Approach to Breast Cancer Detection Using Light," SPIE vol. 1888 pp. 500-510, (1993). Changes in the local optical properties of tissue can be measured by intersecting an acoustic field with a radiation field (see FIG. 4). Specific implementations can provide three dimensional information.
A problem addressed in co-pending application Ser. No. 08/480,760, by the present inventors, is that human skin has an index of refraction for non-ionizing radiation significantly different from that of air. In addition, human skin is not smooth on a microscopic scale and may also exhibit irregularities on a macroscopic scale. In cases where a transparent compression plate used to flatten the breast at the entrance and/or exit points of the optical radiation beam makes poor optical contact with the skin, or when a compression plate is not used at all (see FIG. 9) then the incident radiation and the exit radiation will be partially reflected and experience additional scattering at the skin surface. A coupling material (such as an appropriate gel or liquid) can be used to reduce the index of refraction mismatch between the skin and the adjacent medium (such as air or a compression plate). See co-pending application Ser. No. 08/480,760, by the present inventors.
In addition, breasts are non-homogenous objects which lack uniform physical dimensions. The thickness of breast tissue over a region to be imaged may not be consistent. If time-resolved optical imaging techniques are employed (e.g., used in heterodyne detection, TOF, holography, etc.), then the optical flight time between source and detector (typically a fixed distance apart) depends not only on the types of tissue encountered as radiation passes through the breast, but also on the total thickness of tissue the light must traverse. A compression plate (enabling shaping a region of the breast) and/or a coupling material can be used to reduce variations in path length.
A further improvement may be the use of optically absorptive materials on parts of the compression plate(s) in order to reduce the level of scatter which could ultimately reach the detector by re-entering the breast. Materials can also be selected on the basis of their acoustic properties if appropriate.
There may be disadvantages associated with imaging from a single direction, which has traditionally produced the familiar projection image common in x-ray radiography. Computed tomography (CT) enhances the available data set by acquiring projections from many directions, permitting three dimensional information to be synthesized from two dimensional data sets. By coupling a collimated beam into the breast (preferably with compression plates or modified compression plates) over a range of angles, angled transmission and backscattered images can be acquired. The data from these multiple projections can be combined to synthesize a three dimensional image (similar to "tomosynthesis" which is practiced in x-ray radiography). If the tissue volume of interest contains contrast-enhancing materials or materials which can be detected through emission fluorescence or Raman scattering (see, J. Wu, et al., Applied Optics Vol. 34, p. 3425-3430 (1995)) or Doppler effects, then the use of multiple collimated angled beams may improve localization capabilities. The angled transmission and backscatter measurements can also be made in conjunction with an intersecting acoustic field.
Angled incident beams or adjacent parallel beams have been used to provide a measurement for an estimated scatter correction in the collimated output. The radiation scattered in the appropriate direction can also be used as a virtual collimated beam. This virtual beam appears to originate from a tissue volume different from the externally incident collimated beam. Multiple projections can also be acquired using the virtual collimated beam data. Virtual collimated beam information can also be used to enhance the tomosynthesis image based on the collimated optical beam data.
The advantages of using compression plates were described earlier. In some instances it may be beneficial to modify the design by providing an open area in the compression plate where only air or a coupling fluid/gel are in contact with the skin surface of the area to be imaged. The region of the breast in the open area will still be of a relatively uniform thickness. Such an open area can be implemented for one or both plates. An open area compression plate can be used with the previously discussed optical and acousto-optical imaging techniques, and also to couple acoustic radiation into the compressed breast. The effect of this acoustic radiation beam (which can be tilted manually or electronically if desired) on tissue can be observed with an intersecting optical beam (which can also be tilted if desired). In addition, acoustic imaging techniques benefit from compression since tissue thickness over a region can be reduced significantly. The concepts of multiple transmitted and backscattered collimated angled optical beams, virtual collimated optical beams, and tomosynthesis can now be extended to include acoustic radiation beams.
Compression plates, with or without an open area, reduce the effective volume of tissue which needs to be evaluated and therefor can be useful for uncollimated sources and receivers as well as the collimated designs we have discussed. Thus, compression plates can be used to an advantage for conventional uncollimated diffuse and diffusive wave optical or acoustic (or acousto-optic) imaging techniques.
A new device and/or mode of imaging breasts using optical, acousto-optical, and acoustic non-ionizing radiation imaging techniques is needed to improve image quality and tissue characterization accuracy. Particular problems which need to be addressed are the thickness of a typical breast for optical, acousto- optical, and acoustic data acquisition, the need for improved radiation coupling into and out of the breast, the desirability of sampling volumes of uniform thickness, the need to enhance the information content of detected radiation, and the desirability of sampling a tissue volume from more than one direction.
Prior devices and methods do not address these concerns.